The present invention relates generally to Magnetic Resonance Imaging (MRI) systems, and more particularly, to a system for generating a highly uniform static magnetic field.
Magnetic Resonance Imaging (MRI) systems are generally in either a cylindrical or an open architecture configuration. Both the cylindrical and open architecture configurations include a superconducting magnet that generates a temporally constant primary magnetic field. The superconducting magnet resides within a cryostat, which cools the superconducting magnet and maintains the operating temperature thereof. The temperature of the superconductor is maintained at approximately 4–10° K., for low temperature superconductors, and at approximately 20–80° K. for high temperature superconductors. The cryostat is typically contained within several thermal shields.
The superconducting magnet is used in conjunction with a first set of magnetic gradient coils, which are sequentially pulsed to create a sequence of controlled gradients in the main magnetic field during a MRI data gathering sequence. The superconducting magnet and the magnetic gradient coil assembly have a radio frequency (RF) coil on an inner circumferential side of the magnetic gradient coil assembly. The controlled sequential gradients are effectuated throughout a patient imaging volume of a patient bore, which is coupled to one or more RF coils or antennae and an RF shield. The RF coils and the RF shield are typically located between the magnetic gradient coil assembly and the patient bore.
As a part of a typical MRI, RF signals of suitable frequencies are transmitted into the patient bore. Nuclear magnetic resonance (nMR) responsive RF signals are then received from the patient via the RF coils. Information encoded within the frequency and phase parameters of the received RF signals, by the use of a RF circuit, is processed to form visual images. These visual images represent the distribution of nMR nuclei within a cross-section or volume of the patient.
When the gradient coils are electrically pulsed, the resulting time changing magnetic flux in any of the electrically conducting cylinders surrounding the gradient coils induces eddy currents. These eddy currents in turn produce their own magnetic fields, which degrade the quality of the desired gradient field in space and time. A second set of gradient coils, sometimes referred to as gradient shield coils, are typically incorporated between the cryostat and the first set of gradient coils to compensate for the aggressive pulse sequences which are routinely used in MRI imaging. The gradient shield coils reduce the amount of mutual inductance between the conducting members, such as the thermal shields, and the first set of gradient coils, which in turn reduces the amount of generated eddy currents.
Also, MR imaging requires a highly uniform magnetic field to generate good quality images. To increase imaging quality it is desirable to increase field strength of the magnet field. By increasing magnetic field strength, the stray MR field increases. Thus, a field strength limitation arises when maintaining the stray MR field below a specified level to minimize or prevent individuals outside of the MRI room from being effected by the field. In order to minimize the field, the MRI room may be shielded from the surrounding environment, however the shielding can be costly and impractical.
For example, a typical 0.5 Tesla open superconducting magnet has approximately 100,000 cubic centimeters of superconductor. This superconducting magnet has a coil-to-coil gap of approximately 60 cm, which yields a room temperature gap of approximately 50 cm for physician access. It is known that a 55 cm room temperature gap is the minimum acceptable for interventional procedures. Thus, prior MRI system designs do not satisfy the minimal acceptable gap requirement, the minimal stray field requirement, and provide the desired static magnetic field uniformity.
It is infeasible to build a magnet of conventional open magnet design that generates a highly uniform magnetic field and has significantly improved physician access for interventional procedures using the known methods of system design. A superconducting magnet of a conventional design that would satisfy the stated design constraints has approximately 392,000 cubic centimeters of superconductor and a corresponding increase in magnetic forces, which is clearly infeasible to build. Similar infeasibilities exist for cylindrical MRI system designs, especially ones of a “Short”or more compact design.
Thus, there exists a need for an improved MRI system that provides increased magnetic field uniformity, minimizes generation of eddy currents and stray fields, and provides increased MRI system design configuration flexibility.